Thermal coagulation therapy may be used for the treatment of localized diseased tissue, e.g., tumors, in a diseased organ or body. Generally, a target volume of tissue is sufficiently heated to achieve a therapeutic effect, such as thermal coagulation. Tissue thermal coagulation depends on a number of factors, and temperatures in the range of 55-60° C. are generally considered sufficient to provide enough energy to cause such coagulation. Cell death results from heating to these temperatures, and a region of irreversible thermal damage can be observed with imaging following the treatment. In addition, heating can be produced from minimally-invasive applicators, eliminating the need for open surgery, and potentially reducing recovery time and morbidity for patients. This approach has been used with some success in the treatment of isolated primary liver cancers and colorectal metastases for patients otherwise ineligible for surgery.
Interstitial thermal therapy is currently practiced by inserting heating applicators directly into a target site within an organ. Several energy sources have been integrated into interstitial heating applicators, including lasers, ultrasound, microwave, and radiofrequency energy. Preferably, interstitial thermal therapy delivers sufficient thermal energy to coagulate an entire target volume, while avoiding undesirable thermal damage to adjacent normal tissues. This strategy is referred to as “conformal thermal therapy.” One limitation of present interstitial thermal therapy technology is the inability to control or adjust the three-dimensional pattern of energy deposition dynamically during a treatment. Most current applicators act as point or line sources of energy resulting in patterns of energy deposition in tissue which are highly symmetrical about the active portion of the applicator. This makes it difficult to treat targets with complex geometry accurately, and does not take full advantage of the imaging information available with imaging technology such as magnetic resonance imaging (MRI).
One application of interstitial heating is transurethral prostate thermal therapy, which selectively coagulates diseased prostate tissue using a device located within the prostatic urethra, and minimizes harm to adjacent normal tissues such as the rectal and bladder walls. Disease targets include prostate cancer and benign prostatic hyperplasia (BPH). It is difficult for current transurethral thermal therapy technologies to produce a thermal treatment (cell death) pattern that conforms accurately to the geometry of the prostate gland or to targeted sub-regions of the prostate gland.
In conformal prostate thermal therapy applications, it is often desirable to implement some form of quantitative temperature monitoring for feedback during treatment to ensure accurate delivery of energy to the prostate gland. Temperature monitoring of treated (or heated) tissue regions can be accomplished in several ways. These include direct measurements as well as indirect measurements of the spatial and/or temporal thermal field in the treatment region.
Magnetic resonance imaging (“MRI”) has been used to measure spatial heating patterns non-invasively in tissue. Several MRI techniques are available to measure the temperature distribution in tissue. These techniques are mainly possible because of the temperature dependence of various nuclear magnetic resonance (“NMR”) biophysical parameters such as T1, T2, diffusion, magnetization, and proton resonant frequency. The most commonly used technique for measuring temperature in MRI-guided thermal therapy is the proton resonant frequency (“PRF”) shift technique, which exploits the direct proportionality between the resonant frequency of water protons and temperature. A common technique to measure this effect employs the subtraction of a baseline phase image obtained prior to heating from a phase image obtained during heating to measure the change in phase resulting from local temperature elevations. The change in phase can then be related to the change in temperature through the expression,
      Δ    ⁢                  ⁢          T      ⁡              (                  x          ,          y                )              =            ΔΦ      t              α      ·      γ      ·              B        o            ·      TE      
ΔT is the temperature change between two images, ΔΦt is the phase change due to temperature differences between the same two images, α is the proton resonant frequency shift coefficient (typically −0.01 ppm/° C.), γ is the gyromagnetic ratio, Bo is the strength of the main magnetic field (T), and TE is the echo time of the imaging sequence used to acquire the two images.
In performing thermal treatments as described above, it is usually preferable to avoid damage to normal (non-diseased) tissue due to heating of the normal tissue in the vicinity of the diseased tissue. This is of special concern for normal tissues proximal to or in the vicinity of the treatment area (sometimes called a target volume or treatment zone) where heat is applied to the diseased tissue. In the example of thermal treatment of the prostate, ultrasound, electromagnetic, RF, microwave, and other sources of heat can lead to heating of normal (non-diseased) tissue surrounding the prostate or adjacent thereto, for example the tissue of the rectum or rectal walls. Excessive heating of this normal tissue could cause unwanted damage to the normal tissue, which could contribute to patient morbidity.
Various attempts to prevent over-heating of tissue outside the targeted diseased tissue have included attempts to confine the volume within which the thermal therapy is applied so that thermal effects are reduced beyond the localized volume being treated. This solution can result in slower or less effective treatment, as treatment of an extended region in space would require application of many such localized treatments to a small treatment zone to avoid spreading the thermal energy to normal tissues outside the treatment zone. Other attempts to account for the heating of normal tissues outside a target volume use the cooling effects of time so that short duration pulses of heat are applied to a target area in order to avoid over-heating of various locations and to enable conduction or other heat transfer mechanisms to keep the temperature of normal tissues in check. Yet other attempts to counter the effect above include using the heat transfer capabilities of perfused tissue having blood flowing therethrough to carry away certain doses of heat applied to the tissue.
Therefore, it remains needed or useful to develop techniques and treatment systems to heat the diseased tissues in a targeted treatment volume sufficiently while at the same time avoiding over-heating of the proximal tissues and organs. In particular, the present embodiments and concepts will illustrate to one skilled in the art methods and apparatus for achieving an effective, robust, efficient thermal treatment of diseased tissue simultaneously with no, little, or reduced risk of damage to nearby normal tissues and organs under given conditions.